Generally, laser photocoagulation has been the standard treatment of a number of ocular and non-ocular disorders. In particular, ocular disorders such as proliferative diabetic retinopathy (PDR), diabetic macular edema (DME), choroidal neovascularization (CNV), retinal tears, and retinal vein occlusion (RVO) can be successfully managed using laser photocoagulation. Additionally, laser photocoagulation has been successfully used for treatment of cutaneous vascular and pigmented lesions. Retinal laser photocoagulation settings utilized for ocular treatments, indications, and outcomes of the treatments have been well evaluated by multiple national trials such as the diabetic retinopathy study (DRS), the early treatment of diabetic retinopathy study (ETDRS), and the macula photocoagulation study (MPS). The vast majority of these studies were conducted utilizing visible lasers, particularly in the green range produced by Argon or Krypton gas lasers. All of these studies provided evidence for the beneficial use of lasers in preserving vision and treating ocular disorders.
Generally, photocoagulation involves heating up the tissues to produce irreversible cellular damage to decrease the metabolic demand of the tissues. This is achieved by irradiating the tissues with a high fluence of laser radiation, that in turn is absorbed by the by tissue chromophores, most significantly melanin. The chromophores build up heat that diffuses to surrounding tissues to cause photocoagulation of the adjacent tissues. All of the above laser treatments involve longer pulses of a continuous wave (CW) laser to photo-coagulate the tissues, typically for 50 milliseconds or longer. A progressive rise in tissue temperature in laser photocoagulation is caused when the rate of tissue temperature rise exceeds that of heat dissipation, which does not allow for cooling down of tissues.
Ophthalmic laser treatments have been shown to result in therapeutic effects without causing irreversible photocoagulative damage to the ocular tissues. This non-photo-coagulative therapeutic effect can be obtained utilizing short-pulsed lasers as discussed in the early work by Pankratov, “Pulsed delivery of laser energy in experimental thermal retinal photocoagulation”, Proc. SPIE, V1202, pp. 205-213, 1990, wherein it is suggested that when the laser was pulsed in the frequency range of 1-20 kHz, the most influential factor determining the size of the laser burn was the duty cycle. Pankratov reported that the smaller the duty cycle, the smaller the lesion. The short pulse laser treatments have demonstrated beneficial effects while minimizing deep choroidal damage by using a pulse train with a low duty cycle to confine the thermal effects to the cells with the chromophores.
Q-switched lasers are capable of producing very short laser pulses, typically in the range of 3-10 nanoseconds. Q-switched lasers have been used for selective laser trabeculoplasty (SLT) treatment of the trabecular meshwork for treatment of glaucoma and selective retinal treatment (SRT) for clinically significant diabetic macular edema (CSME) via selectively targeting the melanophores within the trabecular meshwork and the retinal pigment epithelium (RPE) respectively. However, the use of such short pulses with very high max power over a very short period of time results in plasma formation causing photo disruption and mechanical shock wave in the targeted tissues. This process can cause bubbling and mechanical damage of the targeted tissue.
Sub-lethal laser treatments can be performed either with a long, CW, low energy level laser treatment, or via a CW laser with short and controlled pulse widths. Low-level laser therapy (LLLT) using low level laser (LLL) is a non-thermal medical laser treatment modality commonly used with photo-bio-modulation. LLLT has become an increasingly used laser modality. While LLLT has been used mainly for wound healing and pain relief, the medical applications of LLLT have broadened to include diseases such as stroke, myocardial infarction, and degenerative or traumatic brain disorders. For example, in 1967, Endre Mester noticed the ability of the low power Helium-neon (HeNe) laser to increase hair growth and stimulate wound healing in mice (Mester E, Spiry T, Szende B, et al. Effect of laser rays on wound healing. Am J Surg. 1971; 122:532-535.).
There are different protocols for LLLT but the common features include irradiating the tissues with a visible or infrared laser of low fluence of 10 mW/cm2−0.5 W/cm2. LLLT acts by inducing a photochemical reaction in the tissues at a cellular level. LLLT has multiple recognized modes of action, most prominently the ability to reduce inflammation. In particular, LLLT produces an anti-inflammatory effect via lowering levels of prostaglandin E2, prostaglandin-endoperoxide synthase 2, interleukin 1-beta, tumor necrosis factor-alpha, and the cellular influx of neutrophil granulocytes, oxidative stress, edema, and bleeding. The appropriate dose for LLLT has been suggested to be between 0.3 and 19 joules/cm2 (Bjordal, J. M.; Johnson, M. I.; Iversen, V.; Aimbire, F.; Lopes-Martins, R. A. B. (2006). “Low-Level Laser Therapy in Acute Pain: A Systematic Review of Possible Mechanisms of Action and Clinical Effects in Randomized Placebo-Controlled Trials”. Photomedicine and Laser Surgery 24 (2): 158-68. doi:10.1089/pho.2006.24.158).
Alternative to LLLT, there are lasers that generate a pulse train of short pulses (typically 100 to 10,000 microseconds) with higher power during the pulse, but significant off time between pulses (typically 5 to 25% ON duty cycle and 75-95% OFF). This allows the energy to be confined in a small area using the thermal dissipation principle. The thermal confinement results in focal and localized therapeutic effect because the specific absorbing tissue, which is being heated by the pulses to temperatures above the standard photocoagulation threshold, are activated but not irreversibly coagulated; since the heat can dissipate fast enough that no coagulation takes place. This method allows significant treatment without causing full thickness retinal damage and the associated vision loss.
Currently available lasers for retinal laser treatments include visible green, green-yellow, yellow or red lasers that are obtained via diode pumped solid state (DPSS) laser sources. The standard DPSS photocoagulator lasers require continuous and fast sensing of the actual laser output power level of the laser cavity using a photodetector and feedback current control mechanisms because of the fluctuation of the laser output with the change of the temperature of the laser pumping diode, laser gain medium, and frequency doubling crystal.
DPSS lasers, unlike diode lasers, are very susceptible to variations in the output power of the pumping diode. This variation is due to a number of reasons, but two of the reasons are particularly important: 1) The laser gain medium and the second harmonic conversion crystal are both optically non-linear in response to their respective excitation wavelengths, therefore alterations in the original pump diode output power can lead to unexpected effects including unstable performance and increased optical noise; and, 2) The thermo-mechanical stability of the DPSS laser is controlled very carefully to ensure stable optical performance, varying the pump diode power can affect this thermal equilibrium to the detriment of the laser performance and lifetime. Additionally, despite active temperature management of the different components of the DPSS cavity, the temperature control is much slower than the fluctuation of cavity component temperature. Moreover, all DPSS lasers, regardless of the wavelength, have limited modulation speed. Furthermore, DPSS sources when compared to laser diodes are well known for their low power efficiency, high sensitivity to temperature, and the need for active cooling, complexity, larger size, and weight, progressive loss of power, higher cost, and shorter life span.
Recently, short pulsed visible lasers have become available in both green (532 nm) and yellow (577 nm) through a DPSS source controlled with software and/or hardware control loops (see U.S. Pat. No. 7,771,417). The software control loop, in response to the signal from the photodetector, alters drive current to the visible laser diode within about a millisecond, and the hardware control loop, in response to the signal from the photodetector, controls timing of the train of pulses to within a microsecond. The DPSS controller provides instructions for the output of the train of pulses with on times of 25 microseconds to 10 milliseconds per pulse, such that the train of pulses is sufficient for photoactivation of a therapeutic healing response in tissue at a target site and off times of 31.67 to 100,000 microseconds, such that the train of pulses is insufficient to induce traditional photocoagulation of the tissue at the target site. Additionally, the photodetector couples the controller to the pulsed output so that a power of the pulsed output is within less than 10% of a desired power level.
Additionally, the present invention assumes that the therapeutic effect of the non-thermal lasers is related to the sub-lethal tissue heating from pre-treatment temperature to post-treatment temperature, which triggers the cascade of processes responsible for tissue healing. Despite the lack of lethal damage caused by excessive tissue heating, the heat dissipated from the tissue chromophores causes intracellular warming that can be gradually additive at the cellular level, but not enough to raise the cellular temperature above the critical level of photocoagulation of the cellular proteins. More laser energy is required initially to raise the tissue temperature from the baseline temperature to the desired subcritical temperature. Once the cellular temperature is within the required temperature range, less laser energy is required to maintain temperature to avoid reaching the critical temperature leading to irreversible tissue damage. The currently available non-photocoagulative laser treatments allow laser pulses with fixed energy, fluence, frequency and duty cycle during the on time of laser pulses. Modulating the delivered laser energy to allow for variable laser energy levels per pulse allows optimization of the therapeutic
An alternative method of creating a pulsing a continuous wave (CW) beam is through optical chopping. Traditionally, optical chopping usually involves a rotating disc punctuated with radially arranged holes or slits. As the disc rotates at a certain speed, it periodically blocks the beam and thus pulses the laser at a fixed frequency and fixed duty cycle. The optical chopper is typically driven by an electronic controller which allows for precise control of the rotational frequency of the disc. Because optical choppers are mechanical in nature, the maximum frequency is limited to the high kHz domain, due mainly to air resistance restricting the speed of the optical chopper. The diameter of the laser beam also places some constraints on the size of the slits, and therefore the maximum pulse frequency is lower for larger diameter beams. Optical chopping allows consistent and clean pulse shape while the laser source is running at CW conditions. However, the main limiting factor for pulsing the CW laser for selective tissue targeting, is the inability to change the duty cycle without physically adjusting the rotating disc.